Medical implants can be manufactured from alloys, ceramics or both degradable and stable composites. The choice of implant material selection is always a combination of material property requirements, the type of fixation needed, knowledge and skills of the physician, patient's needs and expectations and sometimes a compromise has to be done between available materials and the requirements of the healing process and the quality of life after trauma, fixation etc. In general, the lack of suitable materials in the market restricts the development and design of certain types of implantable devices.
Traditionally alloys have been used to make bone pins, screws and plates and indeed, for certain applications they still are well suited for carrying external load. However, bone resorption may often be seen due to the strength and stiffness of the alloy compared with the bone. In addition to this hardness problem, another disadvantage is the lack of material degradability in vivo. In order to avoid the bone resorption after the healing process, a second surgery is required to remove the implant, which always causes an additional risk and added morbidity for the patient, occupies the availability of clinic and increases the overall costs (Bradley et. al. Effects of flexural rigidity of plates on bone healing. J Bone Joint Surg 1979; 61A:866-72.).
Biostable polymers and their composites e.g. based on polymethacrylate, ultra high molecular weight polyethylene (UHMWPE), polytetrafluoroethylene (PTFE), polyetheretherketone (PEEK), polysiloxane and acrylic polymers are known in the literature (S. Dumitriu, Polymeric Biomaterials 2nd ed., CRC Press, 2001), and polymer composites have been used to manufacture medical implants. However, they are neither bioactive nor resorbable and thus will not be replaced by natural bone. Although being weaker than the alloy implants they still suffer similar problems than alloys and may require a second surgery for replacing or removing the implant at some point of the lifetime of the implant.
The biological and mechanical properties of bone result from its microstructural features. Bone is a composite material made up of organic and inorganic components, where the inorganic or mineral phase represents 60-70% of the total dry bone weight. The organic phase is a viscous gel-like material comprised primarily of collagen while the mineral component consists of a crystalline form of calcium phosphate containing carbonate ions, small amounts of sodium, magnesium, hydrogenophosphate ions and other trace elements.
Various bioactive glass compositions are known in the field. They are able to bond to bone and soft tissue, and they may be used for stimulating tissue or bone growth in a mammalian body. Bioactive glass also typically guides the formation of new tissue, which grows within said glass. When bioactive glasses come into contact with a physiological environment, a layer of silica gel is formed on the surface of the glass. Following this reaction, calcium phosphate is deposited to this layer and finally crystallized to a hydroxyl-carbonate apatite. Due to this hydroxyl-carbonate apatite layer the resorption of the bioactive glass is slowed down when inserted into a mammalian body. For decades, bioactive glasses have been investigated as bone filling materials that can bond with bone, even chemically. Recent discoveries of the superior qualities of bioactive glasses have made the materials far more interesting for these applications. Certain bioactive glasses have been commercially sold under the trade names of e.g. BonAlive®, Novabonea and Biogran®. Bioactive glasses have been used in different forms for medical applications, such as granules and plates for orthopaedic and cranio-maxillofacial bone cavity filling and bone reconstruction. Certain bioactive glass formulations have been disclosed in the prior art, e.g. publications EP 802 890 and EP 1 405 647. Some compositions of bioactive glasses are known to have antimicrobial effects, see for example publications U.S. Pat. No. 6,190,643 and U.S. Pat. No. 6,342,207.
Other types of resorbable glass compositions are also known in the field. Resorbable glasses are not necessarily bioactive, i.e. they do not form a hydroxyl-carbonate apatite layer on the glass surface. Resorbable glass compositions are used in the glass fiber industry to resolve the problem of glass fibers ending up e.g. in lungs during installation of glass fiber insulation. Disappearance of the fibers is preferably relatively fast, so that no detrimental effects are caused to the body. One resorbable glass composition is disclosed in document EP 412 878. The fibers are degraded under 32 days. Such degradation rate is, however, too fast for most medical applications, for example for screws or pins for fixing bone defects or fractures.
Documents EP 915 812 and EP 1 484 292 disclose biosoluble glass composition to improve occupational health and safety. Document WO 03/018496 discloses anti-inflammatory, wound-healing glass powder compositions. Publication U.S. Pat. No. 6,482,444 discloses silver-containing bioactive sol-gel derived glass compositions to be used in implanted materials, for preparation of devices used for in vitro and ex vivo cell culture.
Document EP 802 890 discloses a bioactive glass composition with a large working range. Devitrification problems are circumvented by adding potassium and optionally magnesium to the glass.
One aspect of the fiber glass composition is to prevent neuro and/or cytotoxic effects derived from the fiber glass compositions containing potassium and/or a high local pH raise due to a too fast degradation rate of glass fibers.
Although bioactive glass and glass fibers are being well accepted by the body and have proven to be excellent biomaterials for bone fixation applications, bioactive glass lacks the required mechanical properties for load bearing applications. Indeed, bioactive glass is a hard and brittle material.
Resorbable polymers have been used to develop resorbable implants. The advantage of using resorbable polymers is that the polymers and thus the implant resorbs in the body and non-toxic degradation products will be metabolized by the metabolic system. One disadvantage of using non-reinforced resorbable polymers in implantable devices is the lack of mechanical strength and modulus, especially when compared with cortical bone. Another disadvantage of resorbable polymers is that they are not bioactive on their own. In order to achieve a bioactive bioresorbable polymer device, a bioactive compound or compounds, such as bioactive glass, needs to be added to the device. However, the addition of bioactive glass or other bioactive agents typically reduces the mechanical strength even to a lower level than that of the native polymer.
Self-reinforcing has been used to improve the strength of resorbable polymers and medical devices. Self-reinforcing is a polymer processing technique were the polymer molecules are forced to a certain orientation resulting in improved strength of the product. It has been reported that self-reinforced bioresorbable polymeric composites improve the strength of resorbable devices. Indeed, the composites showed relatively good mechanical properties, such as a bending strength of 360+/−70 MPa and a bending modulus of 12+/−2 GPa (P. Törmälä et al., Clinical Materials, Vol. 10, 1992, pp. 29-34), although the reported modulus values were still below the modulus values of strong cortical bone, the bending modulus of human tibial bone having been measured to be 17.5 GPa (S. M. Snyder and E. Schneider, Journal of Orthopedic Research, Vol. 9, 1991, pp. 422-431). The strength and strength retention of self-reinforced poly-L-lactic acid (SR-PLLA) composite rods were evaluated after intramedullary and subcutaneous implantation in rabbits. The initial bending strength of the SR-PLLA rods was 250-271 MPa. After intramedullary and subcutaneous implantation of 12 weeks the bending strength of the SR-PLLA implants was 100 MPa. (A. Majola et al., Journal of Materials Science: Materials in Medicine, Vol. 3, 1992, pp. 43-47).
In order to improve the mechanical strength of resorbable polymer based devices different types of fiber reinforced resorbable polymer composites have been developed. Poly(glycolic acid) (PGA), poly(lactide-co-glycolide) (PLGA), poly(lactic acid) (PLA) fibers in PLA or PDLA (poly(D-lactic acid)) matrix have been manufactured. The initial strength has been very good, however, the PGA and the PLGA fibers resorbed fast and the high strength was lost. Composites in which reinforcing fiber and the matrix were made of the same chemical composition have shown retention of the strength for longer periods of time. Polymer matrix degradation has been slowed down by increasing the hydrophobicity of the polymer and/or by addition of large quantities of buffering agents. Both techniques interfere with the interaction between phases and may result in weakening of the composite. (Publication WO 2008/067531)
However, Törmälä et al. in publication WO 2006/114483 have developed a composite material containing two reinforcing fibers, one polymeric and one ceramic, in a polymer matrix and reported good initial mechanical results, i.e. a bending strength of 420+/−39 MPa and a bending modulus of 21.5 GPa, which are the same level as for cortical bone. However, they have not reported any in vivo or in vitro hydrolytic behaviour and the prior art teaches that bioabsorbable composites, reinforced with absorbable glass fibers, have a high initial bending modulus but that they rapidly lose their strength and modulus in vitro.
The interaction of the hard brittle mineral phase and the flexible organic matrix gives bone its unique mechanical properties. The development of bone repairing materials or substitutes is typically oriented to combinations of mineral materials i.e. bioresorbable glasses to an organic polymeric matrix in order to generate a composite material exhibiting the toughness and flexibility of the polymer and the strength and hardness of the mineral filler and/or reinforcement. Numerous patents disclose the preparation and composition such a composite material (WO 2006/114483, U.S. Pat. No. 7,270,813, WO 2008/067531, WO 2008/035088).
The ultimate aim for a biomaterial in the field of bone and fracture fixation is that the material should mimic all the properties of bone, be bioactive, osteoconductive and biocompatible. Although the composite materials in prior art have led to the composite materials with attractive characteristics, they are still in need of improvement. At the moment, none of the prior art composites have been shown to possess in vivo mechanical properties comparable to natural bone.
A typical problem of the prior art composites is a poor polymer to reinforcement interface interaction and adhesion. The poor adhesion between the polymeric matrix and the ceramic reinforcement results in early failure at the interface in a physiological environment, and therefore the mechanical properties of the composite degrade too fast. Such degradation usually happens through hydrolysis of the interface. Therefore, improvement of the interfacial bonding (covalent bonding) is a key to the successful application of the biodegradable polymer composites to medical fields.
In the absence of a good interfacial adhesion between the polymer and inorganic reinforcement, transfer of stresses experienced by the load-bearing composite material from the elastic polymer to the stiff reinforcement will not appear. A lack of real covalent bonding/adhesion between the two phases results in early failure of mechanical properties in hydrolytic environment. Coupling agents, such as silanes, find their largest application in the composite industry, the compatibility between the reinforcement and polymer having long been known to be improved by using several types of surface coatings and coupling agents. Typically, any silane that enhances the adhesion of a polymer is often termed a coupling agent, regardless of whether or not a covalent bond is formed.
In the field of biomaterials, similar methods have recently been applied to improve the interface of hydroxyapatite or Bioglass®/polymer composites using coupling agents. However, in most of the cases, these treatments result in significant improvements in the ultimate stiffness of the composite (such as in WO 98/46164), but one major drawback lies in the fact that when the polymer matrix is made of biodegradable polymers they lack real covalent bonding between the reinforcement or filler and the polymer backbone and/or the reactive end-groups due to none-existence of them or low amount of them because of too high molecular weight (molecular weight of over 30 000 g/mol). Attempt to form covalent bonds into a weak polymer backbone typically leads to random chain scission, very low molecular weight fragments, gas evolution, unsaturation and autocatalytic degradation, which will ultimately lead to poor mechanical properties and thermal instability of the composite.
Moreover, similar type of methods have been applied to non-bioresorbable composites (see for example document U.S. Pat. No. 6,399,693). However, these materials are known to be highly resistant to hydrolysis and resorption both in vivo and in vitro. These materials would thus have similar disadvantages as metals and biostable materials, such as bone resorption and stress shielding when used as implant material in medical devices.